Positron emission tomography (PET) is becoming a powerful modality to image cancer and other disease. It is the most accurate non-invasive method for measuring the concentrations of radiolabeled tracers in different locations of the body. PET is capable of imaging and measuring the concentrations of a particular biochemical, which in turn provides important physiological parameters in specific locations or organs. PET is an imaging modality that provides biochemical and physiologic information, whereas CT scans or MRI provides anatomical or structural information (Daghighian F, Sumida R, and Phelps M E.: “PET Imaging: An Overview and Instrumentation” J. Nucl. Med. Tech. 18, 5 (1990)).
The basic principle behind PET is that positrons emitted by positron emitting isotopes find an electron and annihilates to two identical photons that travel in opposite directions. The patient is injected with a positron emitting radio-pharmaceutical, such as F-18 labeled flourodeoxyglucose. This radio-pharmaceutical accumulates in the cancer tissues in amounts greater than other tissues. The patient is surrounded by a ring of detectors that are tuned to detect the annihilation photons of the positron-electron annihilation that occurs in the regions where the radio-pharmaceutical is concentrated. Therefore the positron emission is detected based on the detection of two annihilation photons by gamma ray detectors of the PET scanner. The computer portion of the PET scanner records the location of the two detectors that were hit by such photons within a time window of a few nanoseconds (coincidence time window). This coincidence detection of the annihilation photons is an essential part of the positron emission tomography. The position of the positron source is on the line that connects these two detectors, called the “line of response”. The collection of these lines of response allows tomographic reconstruction of the distribution of the radio-pharmaceuticals in the body of the patient, forming the PET images.
Most of the basic elements of biological materials have positron-emitting isotopes (e.g., C-11, N-13, O-15, F-18, I-124). More than 500 biochemicals have been labeled with these isotopes (e.g., amino acids, fatty acids, sugars, antibodies, drugs, neuroreceptor ligands, nucleoside analogues, etc). PET not only provides distribution images of the tracer, but by repeating PET imaging at different times the kinetics of the tracer can be studied. By using an appropriate model, many important physiological parameters can be measured non-invasively at specific locations inside the body.
One of the problems of the current PET designs is the degradation of the spatial resolution away from the axis due to the penetration of the 511 keV photons (the Depth of Interaction or Parallax Problem). In order to reduce this effect, the diameter of the scintillator ring diameter of the standard PET scanner is taken to be larger than the useful field-of-view (FOV) (e.g. 83 vs. 64 cm, for Siemens' ECAT EXACT scanner; and 15 vs. 12 cm for Siemens' microPET scanner). This extra large diameter reduces the sensitivity and increases the cost due to the need for more detector material and associated electronics compared to a system with a smaller diameter. Another problem experienced by PET scanner that causes blurring of the images is the inter-crystal scatter of the annihilation photons.
The preferred basic element of the novel detector module disclosed herein is a photo-detector referred to as a Solid-State Photomultiplier (SSPM), or Silicon Photomultipliers (SiPM). Introduced in 2002, SSPMs have so far been used mainly in high energy and astrophysics experiments where very high sensitivity light detection is required. Such a device is a large assembly of micro pixel diodes operating in a binary mode. Each detector consists of an array of approximately 600 micropixels connected in parallel. The micropixels act individually as binary photon detectors, in that an interaction with a single photon causes a discharge. Each micropixel “switch” operates independently of the others, and the detector signal is the summed output of all micropixels within a given integration time. When coupled to a scintillator, the SSPM detects the light produced in the scintillator by incident radiation, giving rise to a signal proportional to the energy of the radiation. SSPMs have many advantages over photomultiplier tubes (the current standard for scintillation-based detection of radiation). An important advantage is that the operating voltage for SSPMs is around 50 V, as opposed to the kilovoltages required for PMTs, yielding a clear safety advantage for devices to be used inside the body. SSPMs are also extremely small—a 1×1 mm2 detector performs comparably to a PMT with a 1 cm diameter and 5 cm length. SSPMs have an extremely fast signal rise time (˜40 ps), high gain (˜106), good quantum efficiency at 580 nm (>20%), high stability, and low dark current at room temperature. Buzhan P, et al. Nuc. Inst. Meth. Phys Res A , 504, p 48-52 (2003).